1. Technical Field
The technical field generally relates to magnetic resonance imaging and more specifically to a method for improving the image contrast of a magnetic resonance imaging apparatus.
2. Background
Magnetic resonance imaging (MRI), also called nuclear magnetic resonance imaging (NMR imaging), is a non-destructive method for the analysis of materials and is used extensively in medical imaging. It is generally non-invasive and does not involve ionizing radiation. In general terms, nuclear magnetic moments are excited at specific spin precession frequencies, which are proportional to an external main magnetic field. The radio frequency (RF) signals resulting from the precession of these spins are collected using receiver coils. By manipulating the magnetic fields, an array of signals is provided representing different regions of the volume under study. These signals are combined to produce a volumetric image which depends on the T1 and T2 values as well as the spin density.
In MRI, a body is subjected to a constant main magnetic field. Another magnetic field, in the form of electromagnetic RF pulses, is applied orthogonally to the constant magnetic field. The RF pulses have a particular frequency that is chosen to affect particular nuclei, typically hydrogen, in the body. The RF pulses excite the nuclei, increasing the energy state of the nuclei. After the pulse, the nuclei relax and release RF emissions, corresponding to the RF pulses, which are measured and processed into images for display.
When hydrogen nuclei relax, the frequency that they transmit is positively correlated with the strength of the magnetic field surrounding them. For example, a magnetic field gradient along the z-axis, called the “slice select gradient,” is set up when the RF pulse is applied, and is shut off when the RF pulse is turned off. This gradient causes the hydrogen nuclei at the high end of the gradient, where the magnetic field is stronger, to precess at a high frequency, and those at the low end, where the magnetic field is weaker, to precess at a lower frequency. When the RF pulse in a narrow band is applied, only those nuclei which precess at that particular frequency will be tilted, to later relax and emit a radio transmission. That is, the nuclei “resonate” to that particular frequency. For example, if the magnetic gradient caused hydrogen nuclei to precess at rates from about 10.0 MHz at the low end of the gradient to about 10.05 MHz at the high end, and the gradient were set up such that the high end was located at the patient's head and the low end at the patient's feet, then a 10.0 MHz RF pulse would excite the hydrogen nuclei in a slice near the feet, and a 10.05 MHz pulse would excite the hydrogen nuclei in a slice near the head. When a single “slice” along the z-axis is selected, only the protons in this slice are excited by the specific RF pulse to a higher energy level, to later relax to a lower energy level and emit a radio frequency signal.
The second dimension of the image is extracted with the help of a phase-encoding gradient. Immediately after the RF pulse ceases, all of the nuclei in the activated higher energy level slice are in phase. Left to their own devices, these vectors would relax. In MRI, however, the phase-encoding gradient, in the y-dimension, is briefly applied in order to cause the magnetic vectors of nuclei along different portions of the gradient to have a different phase advance. Typically, the sequence of pulses is repeated to collect all the data necessary to produce an image. As the sequence of pulses is repeated, the strength of the phase-encoding gradient is stepped spirally, linearly, or in another fashion, as the number of repetitions progresses. That is, the phase-encoding gradient may be evenly incremented, where the distance between steps is constant, after each repetition. The number of repetitions of the pulse sequence is determined by the type of image desired and can be any integer, typically from 128 to 1024, although additional phase encoding steps are utilized in specialized imaging sequences.
After the RF pulse, slice select gradient, and phase-encoding gradient have been turned off, the MRI instrument sets up a third magnetic field gradient, along the x-axis, called the “frequency encoding gradient” or “read-out gradient.” This gradient causes the relaxing protons to differentially precess, so that the nuclei near the lower end of the gradient begin to precess at a fast rate, and those at the higher end precess at a faster rate. Thus, when these nuclei relax, the fastest ones. i.e., those which were at the high end of the gradient, will emit the highest frequency RF signals. The frequency encoding gradient is applied only when the RF signals are to be measured.
While the z-axis was used as the slice-select axis in the above example, similarly, either the x-axis or y-axis may be set up as the slice-select axis depending upon the desired image orientation and the anatomical structure of the object of interest being scanned. For example, when a patient is laying supine in the main magnetic field, the x-axis is utilized as the slice-select axis to acquire sagittal images, and the y-axis is utilized as the slice-select axis to acquire coronal images. When the area of interest is at an angle to these three planes, different axes may be used to produce oblique imaging. The gradient axes, x, y, and z, may be chosen at an angle to the orthogonal planes by applying a function, usually sine or cosine, to the y and z axes.
Regardless of the orientation of the selected scan, mathematically, the slice select gradient, phase-encoding gradient, and read-out gradient are orthogonal. The result of the MRI scan in the true domain representation k-space is converted to image display data after a 2D or 3D Fast Fourier transform (FFT). Generally, in a transverse slice, the readout gradient is related to the kx axis and the phase-encoding gradient is related to the ky axis. In 3D MRI, an additional phase-encoding gradient is related to the kz axis to acquire data in a third dimension. When the number of phase-encoding levels is smaller than a binary number, the missing data may be filled with zeros to complete the k-space so that an FFT algorithm can be applied.
In k-space, data is arranged in an inhomogeneous distribution such that the data at the center of a k-space map contains low frequency spatial data, that is, the general spatial shape of an object being scanned. The data at the edges of the k-space map contains high frequency spatial data including the spatial edges and details of the object.
The problems associated with MRI image intensity inhomogeneity are well known. Images that exhibit this phenomenon show gradual, low frequency spatial variation in tissue class intensity. The sources of the problem include receiver coil, transmitter coil and magnetic field variations, uncompensated eddy currents, and patient positioning. Display presentation and automatic computer analysis, including tissue segmentation and classification, become problematic with such images.
The receiver coil may be the primary contributor to intensity variations. The spatial variation of the coil field produces images that have strong signal intensities near the coil surface and decreased intensity distant from the coil. Both conventional circumferential coils and, particularly, surface coil arrays may exhibit this problem. Additionally, transceiver coils may exhibit transmitter field and receiver field intensity variations which contribute to image inhomogeniety.
A simple mathematical model of the measured MR image is given by the following equation:R(x,y)=F(x,y)·I(x,y)Where R(x,y) is the received image, F(x,y) is the multiplicative, inhomogeneous coil(s) field, and I(x,y) is the unadulterated true image data. In this model random noise is ignored. If the coil field were known, the received image could be modified by F−1(x,y) producing the more uniform true image. Numerous methods to estimate the receiver coil field have been proposed. One class' of solutions involves knowledge of the coil geometry and electrical characteristics allowing analytic field modeling using the Biot-Savart law. These methods require knowledge of the patient position and size of the receiver coil and do not account for changing coil characteristics. Also, the flexible nature of coil arrays is problematic. Another class utilizes additional measurements on a uniform phantom to map the coil field. The requirement for identical patient and phantom scanning parameters make these techniques impractical. Other techniques use low resolution images acquired at the time of the patient scan to estimate the coil field, thus increasing the scan time. Post-processing or retrospective methods have been proposed also. Some require manual intervention to achieve good results, which is not desirable. Some assume that a low pass filtered version of the image is a good approximation to the coil field, which is not the case in high contrast areas of images. A number of other post-processing methods use image content to generate an estimate of the intensity field. Thus, a method is desired that compensates for image inhomogeneity regardless of the source while simultaneously enhancing the image contrast.